(a) Field of the Invention
The present invention relates to a high-resolution, modular X-ray detector based on a scintillator plate, a lens array, and image intensifiers having a strong optical gain that are optically coupled to solid state detectors, and the use of Maximum-Likelihood Estimation (MLE) techniques to determine a position and energy of scintillation events.
(b) Brief Description of the Related Art
In the fields of digital radiography (DR) and X-ray computed tomography (CT), flat-panel scintillation detectors are used in which the scintillation light can be read out by an array of photodiodes. These detectors sense the total amount of light produced by the x-ray flux over an extended exposure period. Spatial resolution is often limited by the spread of light over from each scintillation event over multiple photodiodes and/or by the size of the photodiode elements. Scintillation detectors are also used in the related field of nuclear medicine where the goal is to image gamma rays produced by a radioactive pharmaceutical injected into a patient or animal subject.
In particular, some recent gamma-ray detectors used in Single-Photon Emission Computed Tomography (SPECT) of small animals make use of photodetector arrays in the form of CCD (charge-coupled device) cameras. In these gamma-ray detectors, a scintillation event is observed as a cluster of signal spread over multiple pixels of the CCD. A few varieties of such detectors exist and each requires the use of sophisticated low-noise, high-quantum-efficiency CCD to observe the scintillation events. Such detectors typically make use of thin scintillators optically coupled to a CCD imager where charge gain is applied within the CCD pixels. Background art CT detectors use thick scintillation plates, typically in the range of 1 mm to 2 mm, and they are physically segmented into individual pixels, typically having a cross-sectional area of 1 mm×1 mm. The physical segmentation minimizes light spread, but the resolution is strongly limited.
Another system used in small-animal single photon emission computed tomography (SPECT) utilizes a scintillator attached to an electrostatic demagnifying tube (DM) which provides slight gain and an increase in the active imaging area, but light loss in the system requires coupling to an EMCCD via a fiber-optic taper to compensate for the losses. Another CCD-based gamma-ray detector is capable of imaging individual gamma-ray interactions using an efficient optical configuration and a low-noise, high-quantum efficiency, cooled CCD camera. Substantial disadvantages of this system are that it only works with relatively thin scintillators that are less sensitive, and the CCD used for the detection must be configured to use long readout time for reduced noise which greatly reduces the frame rate capability of the system.
Moreover, background X-ray scintillation detectors respond to the total amount of light collected on each pixel during a exposure time which may be very long, around 0.1 s. This causes several problems. First, there is a randomness, called Swank noise, due to the variable amount of light, arising in large part from the random X-ray energy. Second, there is a loss of spatial resolution because the light from one X-ray photon can spread to multiple pixels. Third, X-ray photons of different energy are attenuated differently as they pass through the patient's body, leading to image defects in CT referred to as beam-hardening artifacts. Therefore, with the background art it is not possible to take advantage of the information provided by the X-ray photons for optimal diagnosis.
Despite all of the above mentioned solutions in the field of gamma-ray detection as discussed above, there is a strong need for increasing the read-out frequency of the measured scintillations in order to be able to use similar techniques in DR and CT and other applications with rapid arrival of x-ray or gamma-ray photons. Advances in systems are therefore strongly desired requiring high-speed and highly-sensitive X-ray detectors.